Method to minimize chain scission and monomer generation in processing of poly(l-lactide) stent

ABSTRACT

Methods of fabricating an implantable medical devices such as stents made from biodegradable polymers are disclosed that reduce or minimize chain scission and monomer generation during processing steps. The method includes processing a poly(L-lactide) resin having an number average molecular weight between 150 to 200 kD in an extruder in a molten state. A poly(L-lactide) tube is formed from the processed resin and a stent is fabricated from the tube. The number average molecular weight of the poly(L-lactide) of the stent after sterilization is 70 to 100 kD.

CROSS REFERENCE TO RELATED APPLICATION

This application is a continuation of U.S. patent application Ser. No.13/946,964 filed Jul. 19, 2013 which is a continuation of U.S. patentapplication Ser. No. 12/905,918 filed Oct. 15, 2010, each of which areincorporated by reference herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to fabricating implantable medical devices, suchas stents or stent scaffoldings, made from polymers such aspoly(L-lactide).

2. Description of the State of the Art

This invention relates to various kinds of implantable medical devicesincluding structures made from polymers. Such implantable medicaldevices include, but are not limited to, radially expandable prostheses,such as stents and stent grafts, catheters, and pacemaker leads.

Radially expandable endoprostheses are adapted to be implanted in abodily lumen. An “endoprosthesis” refers to an artificial device that isplaced inside the body. A “lumen” refers to a cavity of a tubular organsuch as a blood vessel. Stents are generally cylindrically shapeddevices, which function to hold open and sometimes expand a segment of ablood vessel or other anatomical lumen such as urinary tracts and bileducts. Stents are often used in the treatment of atheroscleroticstenosis in blood vessels. “Stenosis” refers to a narrowing orconstriction of the diameter of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment.

“Deployment” corresponds to the expanding of the stent within the lumenat the treatment region. Delivery and deployment of a stent areaccomplished by positioning the stent about one end of a catheter,inserting the end of the catheter through the skin into a bodily lumen,advancing the catheter in the bodily lumen to a desired treatmentlocation, expanding the stent at the treatment location, and removingthe catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a constraining member such as aretractable sheath or a sock. When the stent is in a desired bodilylocation, the sheath may be withdrawn which allows the stent toself-expand.

The stent must be able to satisfy a number of requirements such as theradial strength necessary to withstand the structural loads, namelyradial compressive forces, imposed on the stent as it supports the wallsof a vessel. Once expanded, the stent must adequately maintain its sizeand shape throughout its service life despite the various forces thatmay come to bear on it, including the cyclic loading induced by thebeating heart. For example, a radially directed force may tend to causea stent to recoil inward. In addition, the stent must possess sufficientflexibility to allow for crimping, expansion, and cyclic loading.Finally, the stent must be biocompatible so as not to trigger anyadverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. The scaffolding is designed so that the stent can beradially compressed (to allow crimping) and radially expanded (to allowdeployment).

Additionally, a medicated stent may be fabricated by coating the surfaceof either a metallic or polymeric scaffolding with a polymeric carrierthat includes an active or bioactive agent or drug. Polymericscaffolding may also serve as a carrier of an active agent or drug.

Furthermore, it may be desirable for implantable medical devices, suchas stents, to be biodegradable. In many treatment applications, thepresence of a stent in a body may be necessary for a limited period oftime until its intended function of, for example, maintaining vascularpatency and/or drug delivery is accomplished. Therefore, stentsfabricated from biodegradable, bioabsorbable, and/or bioerodiblebioabsorbable polymers can be configured to partially or completelyerode away after the clinical need for them has ended.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method of makinga stent comprising: providing PLLA resin having an Mn between 150 to 200kD; processing the PLLA resin in an extruder in a molten state; forminga PLLA tube from the processed PLLA resin; and fabricating a stent fromthe PLLA tube.

Further embodiments of a method of making a stent comprise: providingPLLA resin; processing the PLLA resin in an extruder in a molten stateat a temperature in the extruder is 195° C. to 210° C.; forming a PLLAtube from the processed PLLA resin; fabricating a stent scaffolding fromthe PLLA tube; and sterilizing the stent scaffolding through exposure toe-beam radiation at a dose of 20-35 kGy, wherein the Mn of thesterilized stent scaffolding is at least 70 kD.

Additional embodiments of the present invention include a method ofmaking a stent comprising: providing PLLA resin having a Mn of 150-200kD; processing the PLLA resin in an extruder in a molten state; forminga PLLA tube from the processed PLLA resin; fabricating a stentscaffolding from the PLLA tube; and sterilizing the stent scaffoldingthrough exposure to e-beam radiation at a dose of 20-35 kGy, wherein theMn of the scaffolding prior to sterilization is 100 to 150 kD, andwherein the Mn of the sterilized stent scaffolding is at least 70 kD.

Other embodiments of the present invention include a method of making astent comprising: providing PLLA resin; processing the PLLA resin in anextruder in a molten state; forming a PLLA tube from the processed PLLAresin; fabricating a stent scaffolding from the PLLA tube; andsterilizing the stent scaffolding through exposure to e-beam radiationat a dose of 22-27 kGy, wherein a % change in Mn of the PLLA resin priorto processing in the extruder to after sterilization is less than 70%,wherein the % change is defined as Mn (the PLLA resin prior toprocessing in the extruder)−Mn (the PLLA scaffolding aftersterilization)]/Mn (the PLLA resin prior to processing in the extruder).

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIG. 2 depicts the number average molecular weight as a function ofe-beam dose for samples of poly(L-lactide) resin with different initialnumber average molecular weights.

DETAILED DESCRIPTION OF THE INVENTION

Embodiments of the present invention relate to methods of fabricatingimplantable medical devices such as stents made from biodegradablepolymers. In such embodiments, the methods are directed to reducing orminimizing chain scission and monomer generation during processingsteps. Embodiments are also directed to a stent fabricated from thesemethods. The stents have a reduced concentration of low molecular weightspecies and monomers which are byproducts of thermal and radiationprocessing steps. The methods and devices described herein are generallyapplicable to any implantable medical device. In particular, the methodscan be applied to tubular implantable medical devices such asself-expandable stents, balloon-expandable stents, stent-grafts, andpacemaker leads.

The embodiments are particularly relevant, for reasons discussed below,to implantable medical devices, such as stents, having a polymericsubstrate or scaffolding, a polymer-based coating, a drug-deliverycoating, or a combination thereof. The fabrication methods are furtherdirected to fabrication of stent scaffoldings are the part of the stentstructure which, on their, own, provide patency to a lumen when thescaffolding is expanded within a vessel.

FIG. 1 depicts a view of an exemplary stent 100. In some embodiments, astent may include a body, backbone, or scaffolding having a pattern ornetwork of interconnecting structural elements 105. Stent 100 may beformed from a tube (not shown). FIG. 1 illustrates features that aretypical to many stent patterns including cylindrical rings 107 connectedby linking elements 110. As mentioned above, the cylindrical rings areload bearing in that they provide radially directed force to support thewalls of a vessel. The linking elements generally function to hold thecylindrical rings together. A stent scaffolding such as stent 100 may befabricated from a tube by forming a pattern in the tube with a techniquesuch as laser cutting or chemical etching.

In some embodiments, the diameter of the polymer tube prior tofabrication of a stent may be between about 0.2 mm and about 5.0 mm, ormore narrowly between about 1 mm and about 4 mm. An exemplary tubediameter of the tube that is laser cut to form the stent scaffolding is3 to 4 mm. The wall thickness a tube formed by extrusion or injectionmolding may be 0.05-2 mm, 0.05-1.05 mm, or 0.08-1.02 mm.

Furthermore, stents and other implantable medical devices have beendesigned for the localized delivery of a therapeutic agent. A medicatedstent may be constructed by coating the device or substrate with acoating material containing a therapeutic agent. The underlyingscaffolding or substrate of the device may also contain a therapeuticagent.

In embodiments of the present invention, an implantable medical devicecan be made partially or completely from a biodegradable, bioabsorbable,bioresorbable, or biostable polymer. A polymer for use in fabricating animplantable medical device can be biostable, bioabsorbable,bioresorbable, biodegradable, or bioerodible. Biostable refers topolymers that are not biodegradable. The terms biodegradable,bioabsorbable, bioresorbable, and bioerodible are used interchangeablyand refer to polymers that are capable of being completely degradedand/or eroded when exposed to bodily fluids such as blood and can begradually resorbed, absorbed, and/or eliminated by the body. Theprocesses of breaking down and absorption of the polymer can be causedby, for example, hydrolysis and metabolic processes. In some treatmentsituations, a degradation time of less than 2 years may be desirable,for example, between 6 and 14 months, or more narrowly, between 8 and 12months. Degradation time refers to the time for complete erosion of adevice from an implant site or in vitro.

The present invention is particularly applicable to poly(L-lactide)(PLLA) and poly(DL-lactic acid) (PDLA). The present invention isgenerally applicable to bioresorbable aliphatic polyesters such aspolyglycolide (PGA), polycaprolactone (PCL), and copolymers thereof andwith PLLA and PDLA, such as poly(glycolide-co-lactide) (PLGA),poly(L-lactide-co-caprolactone), and poly(L-lactide-co-DL-lactide). Ingeneral, the present invention is applicable to any bioresorbablepolyester made by ring-opening polymerization. A stent scaffolding canbe made of 100% of one or a combination of the above polymers.Additionally, the scaffolding can also be drug-free.

The molecular weight of polymer of the final or finished stent productis very important for proper functioning or treatment by the stent.Final or finished product can refer to the stent or stent scaffoldingright after sterilization, any time after sterilization, or right beforedelivery in a human patient. The importance of the molecular weightstems from the dependence of properties and behavior of the stent in thebody on molecular weight, most notably, mechanical properties anddegradation properties.

Both the magnitude of the strength and modulus or stiffness of thepolymer depend on molecular weight, being inversely proportional tomolecular weight. The stent property of radial strength depends on thesematerial properties. “Radial strength” of a stent can be definedgenerally as the inward radial pressure at which a stent experiencesirrecoverable deformation. The ability of the stent to resist periodicinward pressure, for example imposed by vessel walls, is directlyrelated to the polymer modulus. For successful treatment, the stent mustat least initially have sufficient radial strength and stiffness tomaintain patency of a lumen at a increased diameter, that is, support alumen at a given diameter.

Additionally, the time dependence of mechanical properties and stentproperties is also significant and dependent on the molecular weight.The stent must be able to support the lumen at the increased diameterfor a sufficiently long enough time period to allow the vessel walls toremodel and remain at the increased diameter without support. As thepolymer degrades, molecular weight decreases and the strength andmodulus decrease along with the radial strength and radial stiffness ofthe stent. After a critical time period the radial strength of the stentdecreases to the point that it can no longer support the lumen (due tothe decrease in molecular weight which causes a decrease in strength andstiffness). Eventually, at another later critical time period afterimplantation, the stent loses structural integrity, i.e., breaks apart.At still another even later critical time period after implantation, thestent completely resorbs away from the treatment site. These criticaltimes periods each are inversely proportional to the molecular weight ofthe finished stent product. Thus, the molecular weight of the finishedproduct must be sufficiently high so that these critical times are highenough for successful treatment. In particular, the finished productmust have a molecular weight high enough so that the stent maintainsradial strength a time long enough to achieve remodeling sufficient forthe lumen walls to be maintained at a desired diameter.

The minimum target molecular weight (Mn) for a coronary stent made ofPLLA with parameters described below is an Mn between 60-110 kD, 60 and100 kD, 60-90 kD, 70-80. An exemplary stent is made of struts having across-section with thickness and width of 140-160 microns, or morenarrowly, 150 microns. The crystallinity of the PLLA stent is 30-55%, ormore narrowly, 40-50%.

The fabrication of a polymer implantable medical device, such as abioresorbable polymer stent scaffolding, can include several processingsteps. These steps include forming a polymeric construct, such as atube, by melt processing; forming a stent body or scaffolding from thetube by laser machining a stent pattern in the tube; and radiationsterilization of the stent. Melt processing can include extrusion orinjection molding. Additional steps can include radially deforming thetube formed from extrusion to increase radial strength, application of atherapeutic coating on the stent body; and crimping the stent over acatheter or delivery balloon.

The polymer during the melt processing, such as extrusion, is free orsubstantially free of a solvent, such as an organic solvent.Substantially free is less than 0.1%, 0.01%, or less than 0.001% ofsolvent as compared to the polymer. Thus, extrusion, as used herein, isdistinguished from techniques such as gel extrusion or solvent extrusionwhich are known to a person of ordinary skill in the art of polymerprocessing. In these methods, a mixture of a polymer and solvent isconveyed through an extruder and a construct is formed there from,followed by removal of all or most of the solvent.

In general, polymers, such as biodegradable and aliphatic polyesters,are subject to chemical degradation during processing. Chemicaldegradation results in a decrease in molecular weight of polymer due tochemical reactions between polymer chains or other species and polymerchains. The chemical reactions, thus, can result in the reduction ordegradation in molecular weight of the polymer. The chemical degradationcan be due to exposure to high temperature, high forces (e.g., shearforces) or due to exposure to radiation. Melt processing and radiationsterilization, in particular, cause a significant drop in molecularweight of the polymer during stent fabrication. The decrease inmolecular weight can adversely effect mechanical properties and otherproperties of the polymer such as degradation behavior and drug releaseproperties.

The chemical degradation of polymers can arise from several differentchemical reaction mechanisms. For example, heat and radiation causechain scission by free radical reactions and non-free radical reactions.Oxygen can accelerate the free radical reactions. Free radical formationresults in chain scission, resulting in the formation of a series ofbyproducts, such as monomers (e.g., lactide monomers from PLLA), cyclicoligomers, and shorter polymer chains. Essentially, chain scissiongenerates new species (byproducts) with unknown or uncharacterizedendgroups. Both melt processing and radiation sterilization result inthe formation of low molecular weight species which are present in thefinal product.

Poly(L-lacide), for example, has hydroxyl end groups and has the generalformula: H—[OCH(CH3)CO]n-OH, which will be abbreviated as: PLLA-OH.Poly(L-lactide) is subject to thermal degradation at elevatedtemperatures, with significant degradation (measured as weight loss)starting at about 200° C. and increasing at higher temperatures. Thepolymer is subject to chemical degradation by both free radical andnon-free radical mechanisms that result in random chain scission thatgenerates byproducts such as oligomers and lactide monomers. To explainthe presence of lactide at higher temperatures, it has been postulatedthat there is an equilibrium between the lactide monomer and the polymerchain. In addition to lactide, the degradation products also includealdehydes, and other cyclic oligomers.

The effect of stent processing (extrusion, laser machining, andradiation sterilization) on properties of PLLA has been observed. Theresorption time of the stent is lower than the resin material that isfed to the extrusion process. For example, a PLLA resin not subjected tomelt processing typically has a reported resorption times of 2 years(Medical Device Manufacturing & Technology 2005) and 3 years (TheBiomedical Engineering Handbook, Joseph D. Bronzino, Ed. CRC Press inCooperation with IEEE Press, Boca Raton, Fla., 1995). An e-beamsterilized PLLA stent can have a resorption time of 18 months to twoyears. Lancet.com Vol. 373 Mar. 14, 2009, incorporated herein byreference.

In melt processing such as in extrusion, the degree of chemicaldegradation and, thus loss of molecular weight, is inverselyproportional to the extrusion temperature. A polymer tube can be formedthrough melt processing methods such as extrusion or injection molding.In extrusion, a polymer melt is conveyed through an extruder barrel toan exit port. The polymer resin is fed to an extruder barrel near itsproximal end as a melt or in a solid form, for example, as a pellet froma hopper. The polymer in the extruder barrel is heated to temperaturesnear or above the melting temperature (Tm) of the polymer and exposed toshear forces and pressures that are generally far above ambient by arotating extruder screw. Table 1 provides melting temperatures reportedfor several polymers. The polymer melt exits the distal end of theextruder barrel into a die. The die imparts a cylindrical shape to thepolymer melt exiting the die, which is cooled to form a tube.Representative examples of extruders for use with the present inventionmay include single screw extruders, intermeshing co-rotating andcounter-rotating twin-screw extruders, and other multiple screwmasticating extruders.

The polymer cannot be conveyed through the extruder barrel unless theviscosity is sufficiently low. The viscosity of a polymer melt is knownto be inversely proportional to the molecular weight. Thus, the higherthe starting molecular weight of the resin, the higher the extrusiontemperature required to provide a viscosity low enough for processing.

Exemplary processing of PLLA resin can be performed with a ¾″ singlescrew extruder. For a resin with a Mn of about 200 kD, the processingtemperature is 200-210° C. and the residence time is 8-10 min. The tubeis quenched in a room temperature water bath as it exits the die. Theextruder barrel pressure is about 2000 psi. The post-extrusion degree ofcrystallinity is about 10%-15%.

TABLE 1 Melting temperatures and glass transition temperatures ofexemplary polymers. Polymer Melting Point (° C.)¹ Glass Transition Temp(° C.)¹ PGA 225-230¹ 35-40 PLLA 173-178¹ 60-65 PCL 58-63¹ (−65)-(−60)60² PDO N/A (−10)-0    ¹Medical Plastics and Biomaterials Magazine,March 1998. ²Science, Vol. 297 p. 803 (2002)

Typically, a stent is sterilized after crimping and packaging thecrimped stent. Irradiation, either gamma radiation or electron beam(e-beam) radiation, are typically used for terminal sterilization ofmedical devices. Radiation sterilization also causes chemicaldegradation to the polymer in a stent. High-energy radiation, such aselectron beams (e-beam) and gamma radiation, tends to produce ionizationand excitation in polymer molecules. These energy-rich species undergodissociation, subtraction, and addition reactions in a sequence leadingto chemical stability. The stabilization process can occur during,immediately after, or even days, weeks, or months after exposure toradiation which often results in physical and chemical cross-linking orchain scission.

The significant drops (discussed below) in molecular weight caused byextrusion and radiation sterilization provide suggest that the molecularweight of the PLLA resin fed to an extruder should have a molecularweight (Mn) as high as possible. Selecting a high molecular weightshould be the best approach to provide a final product molecular weighthigh enough to maintain radial strength for the requisite time periodtreatment and to address safety concerns.

As an example, the inventors selected a PLLA resin with relatively highinitial Mn of 360 kD for the tubing extrusion process. A processingtemperature of about 225-235° C. is required to sufficiently reduce theviscosity for processing the polymer melt. When such high molecularweight resin is used for extrusion, it is difficult to melt uniformlyand the viscosity is very high even at the high processing temperature.The polymer melt is also subjected to high shear force, and pressure,and with relatively long residence time up to 15 min. Such processingconditions caused significant chain scission and produced significantthermal breakdown by-products, and potentially favor the monomergeneration. The molecular weight dropped from 360 to about 250 kD afterextrusion, while the maximum monomer content generated was up to 4%.After e-beam exposure of 25 kGy, the Mn dropped from about 250 kD toabout 85 kD, which is an acceptable molecular weight for treatment.However, as discussed below even though this scheme produces a finishedgood molecular weight in an acceptable range, it results in a productwith a high concentration of degradation byproducts, which inventorshave found adversely effects resorption of the polymer in vivo, i.e.,the resorption rate is increased by the byproducts. This sometimesdramatically decreases the time the radial strength is maintained at alevel that lumen walls can be supported.

Therefore, as summary of the problem with this approach is as follows.It appears reasonable to select a resin with molecular weight as high aspossible to obtain a final molecular weight that is sufficiently high.However, the higher the molecular weight, the higher the extrusiontemperature required to lower the viscosity. The disadvantage of this isthat the higher temperature increases the amount of unfavorabledegradation byproducts.

Although use of a high molecular weight starting material can provide astent with acceptable properties for safe treatment, the presentinvention represents an improvement over this method for reasonsdiscussed below. Embodiments of the present invention are based on therealization by the inventors that the degradation behavior depends onfactors in addition to the molecular weight of the final product.Specifically, the degradation byproducts generated by melt processingand radiation significantly increases the degradation rate, increasingthe rate that radial strength deteriorates upon implantation. Therefore,the time that a stent can maintain support of a lumen decreases withmore byproducts. These byproducts include low molecular weight oligomersand monomers.

Therefore, selecting a resin molecular weight as high as possible thatrequires a high melt processing temperature that results in morebyproducts that are adverse to the stent properties. To the inventors'knowledge, prior to this realization of the link between such byproductsand stent properties, it has not been recognized that thermaldegradation byproducts can significantly compromise the degradationproperties of a polymer stent, in particular, the time that a stent cansupport a lumen.

Embodiments of the present invention include selecting the molecularweight of the polymer resin, for example PLLA, in a manner that iscontrary to the approach outlined above of selecting the molecularweight as high as possible. In the embodiments of the present invention,a lower molecular weight resin is selected which then allows a lowerprocessing temperature. This lower processing temperature necessarilyresults in less harmful byproducts as compared to the approach discussedabove.

In some embodiments, a relatively low melt processing temperature isselected and a molecular weight of the resin is selected that it has aviscosity at the temperature that is low enough for extrusion. In someembodiments, the selected molecular weight may be the highest molecularweight that can be processed at the selected temperature. In otherwords, a resin with a molecular weight higher than the selectedmolecular weight would have a viscosity too high to process at theselected temperature. In the case of PLLA, the selected temperature maybe 20-30° C. higher than the Tm and Mn of the resin is 150-200 kD.

The inventors have further recognized the use of a lower molecularweight resin (150-200 kD) can provide a finished stent molecular weightthat is the same or slightly lower than the use of a higher molecularweight resin (250-350, 350-400, 400 or higher). This is the case evenwith an initial resin molecular weight that is significantly lower, forexample, less than 50% of a higher molecular weight resin.

This recognition is based on data showing that the degree of chainscission from e-beam exposure on PLLA is inversely proportional with themolecular weight pre-exposure. In other words, the degree of chainscission decreases with increasing molecular weight.

FIG. 2 depicts the Mn as a function of e-beam dose for samples of PLLAresin with different initial Mn. Sample preparation and sterilizationare given below in the section labeled Example of this application. Thesample details are provided in Table 2. As shown in the table, thesamples are categorized by three crystallinity groups and threemolecular weight groups.

TABLE 2 Number Average MW (Mn) of Disk Samples (Dose = 0 kGy) NumberAverage MW - Mn (kDa) Low Medium High Crystallinity Low Group A Group DGroup G (15%)* 122 231 N/A Medium Group B Group E Group H (35%)* 114 247375 High Group C Group F Group I (55%)* 116 185 N/A

As shown in FIG. 2, the change in Mn with e-beam dose is very sensitiveto initial Mn. The initial high Mn samples, 250 kD and above, experiencevery high losses with small changes in e-beam dose (e.g., at 25 kGy, Mnof sample H is only about 35% of initial Mn). The initial medium Mnsamples are less sensitive to dose but still experience relatively largelosses (at 25 kGy Mn of sample F is about 55% of initial Mn). Theinitial low Mn samples are relatively flat (compared to the othergroups) across the range of dose evaluated (0 to 50 kGy). This isespecially true within the possible range of sterile doses to beexperienced by the stents during radiation sterilization (20 to 35 kGy).Starting with a higher initial Mn provides a higher Mn drop, within therange of e-beam doses that the stent could possibly see (20-35 kGy),than starting with a lower initial Mn. The change in molecular weightwith e-beam dose is not effected by initial crystallinity over the rangeof crystallinity from 5-55% for low and medium initial molecular weight.

It is clear from FIG. 2, that the e-beam process acts as an “equalizer”of sorts between starting with high molecular weigh and lower molecularweight. The change in molecular weight being more significant for highmolecular weight materials subjected to e-beam brings the molecularweights for high and low molecular weight starting materials much closerafter e-beam exposure.

As noted above, sterilization is typically the final step of themanufacturing process for a stent. The extrusion step accounts for mostof the loss in molecular weight just prior to sterilization. Lasermachining accounts for another potential loss in molecular weight.Therefore, the sample molecular weights correspond to the molecularweight of the stent after losses due to at least the extrusion and lasermachining steps. To determine the molecular weight of a resin that wouldgive the post-radiation exposure molecular weight at a given dose inFIG. 2, the loss in molecular weight from melt processing and lasermachining at least would have to be taken into account.

The use of a resin with a certain lower range of Mn has severalprocessing advantages for extrusion. This includes that the resin can bereadily melted more uniformly with low viscosity at a lower temperature.Additionally, lower shear force and pressure can be applied with shorterresidence time. These factors result in fewer degradation byproductsduring extrusion.

Additionally, the use of such lower molecular weight resin incombination with lower processing temperatures will result in a finishedgood with a significantly lower concentration of degradation byproducts.With respect to the melt processing at lower temperatures, fewerdegradation byproducts are generated than processing a higher molecularweight polymer at higher temperatures. The relative decrease inmolecular weight during melt processing is expected to be lower than athigher temperatures. For example, an extrusion temperature in the rangeof 195-210° C. as compared to 225 to 235° C. Therefore, there will befewer low molecular weight species with various end groups and monomersgenerated.

Another advantage of using the lower molecular weight resin polymer isthat it results in a lower molecular weight polymer just prior toradiation sterilization. The lower molecular weight construct willundergo significantly less chain scission during e-beam exposure (asshown by FIG. 2) than the higher molecular weight construct. Therefore,both the extrusion and e-beam steps will generate fewer degradationbyproducts and the finished product will have significantly fewerdegradation products than a finished product made from an initial highermolecular weight resin.

As indicated above, these degradation byproducts can significantlycompromise the degradation properties of a polymer stent such as thetime dependence of mechanical properties. The finished good from thelower molecular weight resin will maintain mechanical properties longerand more consistently than a finished good from the high molecularweight starting material.

The PLLA resin for the various embodiments of the present invention maybe obtained commercially. For example, PURAC sells PLLA of variousmolecular weights under the trade name PURASORB®. These include:PURABSORB® PL 18 (inherent viscosity=1.8 dL/g), PURABSORB® PL 24(inherent viscosity=2.4 dL/g), PURABSORB® PL 32 (inherent viscosity=3.2dL/g), and PURABSORB® PL 49 (inherent viscosity=4.9 dL/g).

Certain embodiments of fabricating a stent include providing a PLLAresin. The resin melt is processed in an extruder in a molten state toform a PLLA tube. In some embodiments, the processing temperature in theextruder is less than 210° C., less than 200° C., 195 to 210° C., or 195to 210° C. A stent scaffolding is made from the polymeric tube bycutting a pattern into the tube, for example, by laser cutting. Thepolymeric scaffolding is exposed to e-beam radiation for sterilization.The Mn of the scaffolding after sterilization is 70 to 110 kD, or morenarrowly, 80-90 kD.

In this and other embodiments, the fabrication of the stent can includeadditional processing steps including radial expansion of the tube priorto cutting the pattern, coating the scaffolding with a mixture of drugand polymer, and crimping the stent scaffolding over a delivery balloon.

Additional embodiments of fabricating a stent include providing a PLLAresin with an Mn between 150 to 200 kD. The resin is melt processed inan extruder in a molten state to form a PLLA tube with the processingtemperature in the extruder being 195 to 210° C., or more narrowly, 195to 210° C. A stent scaffolding is made from the polymeric tube bycutting a pattern into the tube, for example, by laser cutting. Thepolymeric scaffolding is exposed to e-beam radiation for sterilization.The Mn of the scaffolding prior to sterilization is 100 to 150 kD, ormore narrowly, 110-130 kD. The Mn of the scaffolding after sterilizationis 70 to 110 kD, or more narrowly, 80-90 kD.

Further embodiments of fabricating a stent include providing a PLLAresin and melt processing in an extruder in a molten state to form aPLLA tube, making a stent scaffolding from the polymeric tube by cuttinga pattern into the tube, for example, by laser cutting, and exposing thepolymeric scaffolding to e-beam radiation for sterilization. In someembodiments, the change in Mn from the PLLA resin prior to meltprocessing to after sterilization (final product) is less than 200%,less than 100%, less than 50%, 50-100%, 100-150%, or 150-200%. The %change in molecular weight is defined by: 100%×[Mn (resin)−Mn (finalproduct)]/Mn (resin). In other embodiments, the % change in Mn of thestent scaffolding immediately prior to sterilization to aftersterilization is less than 40%, 40-60%, or 40-50%, where the % change inMn is defined similar to above.

The methods of the present invention applies to stent scaffoldings madefrom other polymers including, but are not limited to, poly(trimethylenecarbonate), polydioxanone, poly(hydroxy butyrate), poly(butylenesuccinate), polyesteramide, poly(hydroxy butyrate-co-hydroxyvalerate),poly(butylene succinate adipate), poly(hydroxyl alkanoates),poly(hydroxyl butyrate), poly(hydroxyl hexanoate), and poly(hydroxylvalerate).

For the purposes of the present invention, the following terms anddefinitions apply:

The term “molecular weight” can refer to one or more definitions ofmolecular weight.

“Molecular weight” can refer to the molecular weight of individualsegments, blocks, or polymer chains. “Molecular weight” can also referto weight average molecular weight or number average molecular weight oftypes of segments, blocks, or polymer chains. The number averagemolecular weight (Mn) is the common, mean, average of the molecularweights of the individual segments, blocks, or polymer chains. Molecularweight is typical expressed in grams/mole which is referred to as“Daltons.” It is determined by measuring the molecular weight of Npolymer molecules, summing the weights, and dividing by N:

${\overset{\_}{M}}_{n} = \frac{\sum\limits_{i}{N_{i}M_{i}}}{\sum\limits_{i}N_{i}}$

where Ni is the number of polymer molecules with molecular weight Mi.The weight average molecular weight is given by

${\overset{\_}{M}}_{w} = \frac{\sum\limits_{i}{N_{i}M_{i}^{2}}}{\sum\limits_{i}{N_{i}M_{i}}}$

where Ni is the number of molecules of molecular weight Mi.

The “inherent viscosity” (of a polymer) is the ratio of the naturallogarithm of the relative viscosity, ηr, to the mass concentration ofthe polymer, c, i.e. ηinh=(ln ηr)/c, where the relative viscosity (ηr)is the ratio of the viscosity of a polymer solution, η, to the viscosityof the solvent (ηs), ηr=η/ηs.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

Example

Sample Preparation

The data for FIG. 2 was generated from test samples that consisted ofsmall disks, or pieces taken from these disks, which were made from PLLAresin supplied by PURAC. The disks were fabricated using a mold andheated press process:

1) A Carver Press was pre-heated to 420° F.

2) Resin was placed into the cavity of the mold, which was sandwichedbetween aluminum foil to prevent the resin from sticking to the press.

3) The molds were then placed into and held in the Carver Press for 15minutes.

The medium crystallinity group was also sandwiched and heated betweenmetal plates to retain heat for a longer period of time during thecooling process.

4) Cooling times:

-   -   a. Low crystallinity—The disk molds were removed from the press        and allowed to cool in the mold for 5 minutes at room        temperature before the disks were removed from mold.    -   b. Medium crystallinity—The disk molds, and metal plates on top        and bottom of the molds, were removed from the press and allowed        to cool in the sandwich configuration for 50 minutes. The disk        molds were then removed from the metal plates and allowed to        cool in the mold for 5 minutes at room temperature before the        disks were removed from the mold.    -   c. High crystallinity—After the 15 minute heating period the        Carver Press heating element was turned off. The disk molds were        left in the press to cool for 24 hours before removing the disks        from the molds.

Nine groups of samples were fabricated. Refer to Table 3 for the name,PLLA resin, target crystallinity and quantity of disks created for eachgroup.

TABLE 3 Disk Sample Groups. Crystallinity Quantity Group Target of NamePLLA Resin (%) Disks A PURABSORB ® PL 18 15 15 B PURABSORB ® PL 18 35 15C PURABSORB ® PL 18 55 15 D PURABSORB ® PL 32 15 15 E PURABSORB ® PL 3235 15 F PURABSORB ® PL 32 55 15 G PURABSORB ® PL 49 15 15 H PURABSORB ®PL 49 35 15 I PURABSORB ® PL 49 55 15

Sterilization

All samples in a given sterilization group (e.g. 12.5 kGy, 25 kGy and 50kGy) were sterilized at the same time to assure they received the sametarget dose.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1-14. (canceled)
 15. A stent scaffolding crimped over a delivery balloonwithin a package, wherein the stent scaffolding is made of alactide-based polymer, and wherein a number average molecular weight(Mn) of the stent scaffolding prior to sterilizing with e-beam radiationis 100 to 150 kD and the Mn of the stent scaffolding after sterilizingwith e-beam radiation is 70 to 110 kD.
 16. The method of claim 15,wherein a dose of the e-beam exposure is between 20 and 35 kGy.
 17. Themethod of claim 15, wherein the polymer comprises poly(L-lactide). 18.The method of claim 15, wherein the polymer comprisespoly(L-lactide-co-caprolactone).
 19. A stent comprising: a stentscaffolding comprising a bioresorbable aliphatic polyester polymerhaving a number average molecular weight of 70 to 110 kD, wherein thestent scaffolding is made by steps comprising: processing a resin of thepolymer in a molten state, the polymer having an Mn between 150 to 200kD prior to the processing step; forming a tube from the processedresin; fabricating a stent scaffolding from the tube; and sterilizingthe stent scaffolding through exposure to e-beam radiation, wherein theMn of the stent scaffolding prior to sterilizing is 100 to 150 kD. 20.The method of claim 19, wherein the polymer comprises poly(L-lactide).21. The method of claim 19, wherein the polymer comprisespoly(L-lactide-co-caprolactone).
 22. The method of claim 19, wherein thestent scaffolding is crimped over a delivery balloon within a package.